This invention relates generally to methods and apparatus for magnetic resonance imaging. In particular, it relates to an electromagnetically de-coupled sandwiched solenoidal array coil for receiving radio frequency magnetic resonance signals in an MRI apparatus.
Based on principals of nuclear magnetic resonance (NMR), magnetic resonance imaging (MRI) has become a widely accepted medical imaging modalityxe2x80x94having evolved tremendously over the last two decades as an important clinical technique for obtaining visual images of tissues and organ structures within the human body. Basically, clinical MRI relies on the detection of NMR signals from abundant hydrogen protons in the human body. These protons are first subjected to a strong radio frequency (RF) electromagnetic wave excitation pulse. If the frequency of the excitation pulse is properly chosen, the protons receive needed RF energy to make a transition to an excited state. Eventually, the excited protons give up their excess energy via a decay process, commonly known as xe2x80x9crelaxationxe2x80x9d, and return to their original state.
Since the magnetic moment of a proton is a vector quantity, the microscopic behavior of millions of protons considered together is equivalent to the vector sum of the individual magnetic moments of all the protons. For convenience, this sum is typically represented as a single resultant magnetization vector, M0, that is aligned with {right arrow over (B)}0 (the static main magnetic field). The strong RF excitation pulse used in MRI effectively tips this resultant magnetization vector away from alignment with the static main field {right arrow over (B)}0 and causes it to precess before decaying back to an equilibrium alignment with {right arrow over (B)}0. The component of this precessing resultant magnetization vector in a plane perpendicular to {right arrow over (B)}0 induces an RF signal, referred to as the nuclear magnetic resonance (NMR) signal, in an RF xe2x80x9cpick-upxe2x80x9d or xe2x80x9creceivexe2x80x9d coil(s) placed near the body portion containing the excited protons.
During clinical MRI, the magnetic resonance of protons in different tissues within an anatomical region are made distinguishable through the evocation of a magnetic field gradient along each of three mutually orthogonal spatial directionsxe2x80x94the effect of which is to cause protons at different spatial locations to have slightly different nuclear magnetic resonance frequencies. The NMR signals induced in the receive coil can then be processed to reconstruct images of the anatomical structure of interest (i.e., images of the spatial distribution of NMR nuclei which, in many respects, conform to the anatomical structures containing such nuclei).
To obtain the maximum induced signal in a receive coil, the magnetic field of the receive coilxe2x80x94conventionally designated as {right arrow over (B)}1xe2x80x94must be oriented perpendicular to the direction of the static main magnetic field ({right arrow over (B)}0) of the MRI apparatus. For a planar-loop (i.e., a substantially flat loop) type receive coil, that direction is in a direction normal to the plane of the conductive loop(s) of the coil. For a quadrature detection (QD) type coilxe2x80x94which basically consists of two RF receive coils having mutually perpendicularly oriented {right arrow over (B)}1 fieldsxe2x80x94must also have the {right arrow over (B)}1 fields of both of its coils oriented perpendicular to the MRI apparatus static field {right arrow over (B)}0 to obtain a maximum induced signal.
Due to the unique nature of the clinical MRI environment, there are certain design considerations that are particularly relevant toward obtaining maximum performance from an RF receive coil. For example, the NMR signals induced in an RF receive coil during magnetic resonance imaging are nominally on the order of nanovolts in magnitude while the background ambient electrical noise may be of comparable levels or higher. Therefore, a high performance RF receive coil for clinical MRI needs to be electromagnetically xe2x80x9csensitivexe2x80x9d enough to detect the low level NMR signals despite the relatively high levels of background electrical noise. Moreover, other design considerations such as field-of-view, uniformity (i.e., uniformity of the magnetic field generated by the coil) and coil efficiency are also highly relevant to coil performance in the clinical MRI environment X coil uniformity because it can affect image interpretation and coil efficiency because a highly efficient coil allows the same image signal information to be acquired within a shorter time frame.
Theoretical analysis and experimental results have indicated that for many MRI applications using multiple RF receive coils together as a signal receiving array is advantageous for improving coil sensitivity, signal-to-noise ratio and imaging field-of-view. Conventionally, the imaging xe2x80x9cfield-of-viewxe2x80x9d (FOV) for an MRI receive coil is defined as the distance indicated between two points on the coil sensitivity profile (i.e., a graph of coil sensitivity vs. distance profile) where the signal drops to 80% of its peak value. In a typical MRI receive coil array arrangement, instead of using a single large FOV but less sensitive coil that covers the entire imaging volume of interest, multiple small FOV but sensitive coils are distributed as an array over the entire imaging volume. Each individual coil of the array covers a small localized volume and the NMR signals received by each coil are simultaneously acquired through corresponding data acquisition channels. Signals from each of the channels are then appropriately combined and processed to construct an image of the complete volume of interest. Due to this ability to simultaneously acquire a signal from multiple sources (i.e., multiple coils) and since each individual signal channel is provided with its own associated detection circuitry, an array type coil can conceivably operate with high efficiency. However, the simultaneous acquisition of a signal from a plurality of individual receive coils necessitates that each coil function independently, free of interaction or coupling.
As two individual coils having the same resonance frequency are brought in close proximity to each other, the common resonance frequency starts to split into two separate frequencies due to the electromagnetic interaction or xe2x80x9ccouplingxe2x80x9d between the coils. Generally, the closer the coils are brought together, the stronger the interaction and the larger the frequency split. Since an MRI receive coil should have its maximum sensitivity optimized for a particular relatively narrow band of frequencies, the resonance frequency splitting can cause sensitivity degradation when two or more receive coils are closely arranged in an array.
Generally, MRI systems are categorized as either a horizontal field type or vertical field type, based on the direction of the static main magnetic field. In a horizontal field system, the static main magnetic field is typically oriented in a superior-inferior direction relative to a patient laying in a prone/supine position. In a vertical field MRI system, the static main magnetic field is oriented in an interior-posterior direction relative to a prone/supine patient. This difference in main field orientation is important in that it affects the ultimate placement and configuration of an RF receive coil(s) used for diagnosis in such systems. More often than not, a receive coil designed specifically for use in a horizontal field system will not be suitable for similar use in a vertical field system and vice versa.
Consequently, horizontal field MRI systems and vertical field MRI systems typically require radically different RF receive coil configurations to obtain the maximum achievable performance from the coil. For example, a planar-loop type receive coil configuration designed for obtaining images of a the human spine works well in a horizontal field MRI system when placed in posterior contact with the back of a patient in supine position. However, the same coil configuration may not work in a vertical field system because, in that case, the RF magnetic field, {right arrow over (B)}1, of the receive coil (i.e., the direction normal to the plane of the loop) is oriented parallel, rather than being perpendicular, to the direction of the static main magnetic field {right arrow over (B)}0.
Likewise, a coplanar loop type array coil is fairly effective in horizontal-field MRI systems but is rather ineffective in vertical-field MRI systems. In a coplanar loop type array coil, the planar-loop receive coils are arranged in a basically coplanar fashion and distributed over the imaging volume of interest. Each individual receive coil in a coplanar loop type array is typically a relatively small but highly sensitive RF coil that receives NMR signals from a specific small portion of the entire region of interest. A final composite image is constructed by combining signals obtained from each of the individual coils.
Magnetic interaction between adjacent coils can be analyzed in terms of induced current, induced EMF or magnetic flux. For the purpose of this discussion, a magnetic flux representation is most convenient. In this representation, two coils in close proximity are considered to xe2x80x9ccouplexe2x80x9d to one another if one coil induces a net non-zero magnetic flux linkage to the other, and vice versa. Likewise, two coils are considered to be magnetically xe2x80x9cde-coupledxe2x80x9d if one coil induces a net zero magnetic flux linkage to the other. Consequently, by definition, a complete nulling of the magnetic flux linkage between coils in an array effectively xe2x80x9cde-couplesxe2x80x9d the individual coils from one another.
Since it has become known that magnetic coupling between adjacent elemental coils in a coplanar loop type array coil can be effectively nulled by properly overlapping the constituent elemental coils (see, for example, xe2x80x9cThe NMR Phased Arrayxe2x80x9d, P. B. Roemer et al., Magnetic Resonant Medicine, 1990, 16, pp. 192-225), various methods and schemes for overlapping the elemental coils have contributed toward making coplanar array coils practical and popular for use in horizontal-field MRI systems. Unfortunately, the coplanar array coils have not been used successfully in vertical-field MRI systems. Although adapting a coplanar array coil configuration to a vertical-field MRI system has been attempted by others, such array coils typically have poor signalxe2x80x94primarily due to the constraint that the normal to plane the of the constituent coils (i.e., the direction of the {right arrow over (B)}1 field) must be positioned perpendicular to the {right arrow over (B)}0 static field of the MRI apparatus to obtain the maximum induced signal.
In an attempt to address this problem, various modifications to the basic coplanar coil array configuration have been proposed by othersxe2x80x94the more familiar modifications being the so-called xe2x80x9cFIG. 8xe2x80x9d array coil and its variations. Nonetheless, known prior art attempts to utilize a coplanar array coil or its variants in a vertical-field MRI system have resulted in severe limitations in terms of coil sensitivity, imaging depth and uniformity over the desired region of interest.
The inventors of the present invention realized that a solenoidal type coil has many inherent characteristics that make it particularly advantageous for use in vertical-field MRI systems. For example, solenoid type coils inherently have high sensitivity and uniformity. In addition, the cylindrical shape of the coil fits naturally over various parts of the body such as the head, neck and other extremities and for clinical MRI applications in a vertical-field system, the B1 field of a solenoidal type coil when fitted to a patient laying horizontal will be oriented perpendicular to the vertical magnetic fieldxe2x80x94as required for maximizing the induced signal strength.
Although it is has been suggested by others that two solenoidal coils be used in a sandwiched arrangement within an NMR spectrometer to achieve mutual isolation between two coils (see U.S. Pat. No. 4,093,910 to Hill, issued Jun. 6, 1978), the two coils comprising this sandwiched arrangement are not used together to form a single RF receiving xe2x80x9carrayxe2x80x9d type antenna for the purpose of obtaining an increase in antenna sensitivity or FOV to provide improved image quality for MRI. Specifically, in the NMR spectrometer, one coil is used as a control channel RF resonance pick-up coil and the other is used as a sample analysis resonance pick-up coil. Moreover, the spectrometer coils are each separately tuned to be responsive to different resonant frequencies (i.e., one coil is tuned for the sample under analysis and the other coil is tuned for a control nuclei, e.g., 2D or 19F). In contrast, a receive coil or receive array coil designed for MRI applications should have its maximum sensitivity optimized for a single relatively narrow band of frequencies.
Accordingly, one general feature of the present invention is an inherently de-coupled array type receive coil having high efficiency, high sensitivity and good uniformity for use in both horizontal and vertical-field MRI systems. The present invention also provides an array type receive coil that can be used in both horizontal and vertical-field MRI systems to provide an enlarged composite field-of-view (FOV) compared to conventional MRI receive coils. The novel array type receive coil can be used in horizontal or vertical-field MRI systems to provide a capability for allowing selection between different sized FOVs and/or imaging regions. It also provides an array type receive coil that can be used in horizontal or vertical-field MRI systems and which can be readily adapted to fit one or more disparate size portions of the human anatomy.
The present invention provides novel RF receive coil array arrangements for enhancing the magnetic resonance imaging of one or more portions of the human anatomy. In particular, the present invention provides an inherently magnetically de-coupled array coil for use in magnetic resonance imaging (MRI) systems. The basic array coil structure of the present invention provides an enhanced signal-to-noise ratio for improved image quality and a selectable field-of-view (FOV) in both horizontal and vertical-field MRI environments. Specifically, the present invention utilizes a field-bucking xe2x80x9csandwichedxe2x80x9d array coil arrangement that precludes magnetic coupling between constituent coils of the array while providing increased sensitivity and uniformity over conventional MRI receive coil arrangements.
In its most basic configuration, the sandwiched array coil of the present invention is an array of radio frequency (RF) sensitive coils comprising at least two solenoidal type RF receive coils arranged such that one coil antenna is coaxially surrounded or xe2x80x9csandwichedxe2x80x9d between two axially separated solenoidal sections of the second coil, with conductor winding directions being opposite in each of the separated sections of the second coil. Because of this xe2x80x9csandwichedxe2x80x9d structural configuration and the peculiar opposite (field-bucking) winding arrangement, the two RF coils of the array are inherently magnetically xe2x80x9cde-coupledxe2x80x9d from one another.
Accordingly, one embodiment of the present invention is a basic sandwiched solenoidal array coil (SSAC) comprising two same-diameter coaxial solenoidal RF coils, each coil having one or more conductive windings, including an inner coil and an outer coil surrounding the inner coil, the outer coil having separated sections with the conductor winding directions being opposite in each of the separated sections so as to create a magnetic field bucking arrangement at the position of the inner coil.
Another embodiment of the present invention is a sandwiched solenoidal array coil (SSAC) having two non-coaxial, unequal-diameter receive coils adaptable to fit portions of the human anatomy. An arrangement for using the SSAC for breast imaging in a vertical-field MRI systems for an example non-coaxial, unequal-diameter sandwiched array coil embodiment is disclosed.
In a third embodiment of the present invention, a sandwiched solenoidal array coil (SSAC) is formed wherein the constituent gradient-field coil of the array has solenoidal sections of unequal diameter and a correspondingly different number of conductive windings in each section.
In a fourth embodiment of the present invention, a sandwiched solenoidal array coil (SSAC) having constituent coils of same or different diameters is adapted for imaging the female breast in a horizontal-field MRI apparatus.
In a fifth embodiment of the present invention, a multiple sandwiched solenoidal array coil (SSAC) apparatus is provided for obtaining region-selectable images of the entire human torso.
In a sixth embodiment of the present invention, a quadrature detection type sandwiched solenoidal array coil is provided for which different data channel acquisition arrangements are shown.